1. Field of the Invention
The invention relates to a measurement device for intracardial acquisition of blood oxygen saturation for frequency control of a heart pacemaker having a measuring probe which contains a light transmitter and a light receiver.
2. Description of the Prior Art
A blood saturation measuring device is described in U.S. Pat. No. 4,399,820 wherein the light receiver receives the light emitted by the light transmitter and reflected by the blood, and wherein the light transmitter and the light receiver are connected to an evaluation circuit via two common lines, this evaluation circuit charging the measuring probe with a constant current or with a constant voltage.
Such a measuring device is schematically shown in FIG. 1. The measuring probe contains a combination composed of a light-emitting diode 6 and a phototransistor 7 which are connected parallel such that the forward current through the light-emitting diode 6 is additively superimposed with the current through the phototransistor caused by the light action. The resistor 8 represents the resistance of the lead. Given drive of the measuring probe by the evaluation circuit A with a constant current or with a constant voltage, the light reflected by the blood dependent on its oxygen saturation triggers a current flow in the phototransistor 7 which effects a change of current or voltage at the measuring probe. The voltage current change generated by the light reflection is identified in an evaluation circuit by comparing the measured signal given a driven light-emitting diode and phototransistor to a reference signal.
FIG. 2 shows a current-voltage characteristic of the measuring probe of FIG. 1. The curve without current through the photo transistor 7, i.e., without reflection, is referenced I. The curve I' then arises with reflection.
The specific sensitivity E of the probe as a function of the blood oxygen saturation SO.sub.2 is an important criterion for the measurement. Given operation with constant current, this shall be referred to below as E.sub.i and shall be referred to as E.sub.u given operation with constant voltage. Thus defined (with units identified in brackets) are:
given constant probe current: i.sub.s =i.sub.k ##EQU1## given constant probe voltage: U.sub.S =U.sub.K ##EQU2## wherein, in accord with FIG. 2,: u.sub.s [mV]=Voltage at the probe (Sensor+Lead) PA1 i.sub.s [mA]=Current through the probe (Sensor+Lead) ##EQU3## Proportion of oxygenated hemoglobin (Hb) in the overall hemoglobin (Hb+HbO.sub.2), i.e.: blood oxygen saturation PA1 u.sub.K '[V]=u.sub.K -u.sub.D =Constant voltage, whereby u.sub.D is the forward voltage of the light-emitting diode 6. PA1 R.sub.L [.OMEGA.]=Lead resistance to the sensor. PA1 K1/100=.multidot.K[%]=Optical coupling factor, dependent on the type of photo element, on the sensor geometry, deposits on the sensor and on the oxygen saturation. ##EQU6## Change of the coupling factor dependent on S.sub.O2
When the specific sensitivity is calculated in accord with FIG. 2, then ##EQU4## whereby: R.sub.D [.OMEGA.]=Series resistance of the light-emitting diode 6. ##EQU5## Slope of the current-voltage characteristic of the light-emitting diode 6 or, respectively, of the sensor given K=0.
It may be seen from equations (3) and (4) that the sensitivity of the probe in both operating conditions decreases with an increasing coupling factor K, a behavior characteristic only for this two-pole embodiment of the reflection probe. As long as k&lt;10%, i.e. K&lt;0.1, the measuring behavior of the probe is unproblematical. Fundamentally, however, an optimally high coupling factor is desired in order to save current and in order to achieve a high SO.sub.2 sensitivity since, of course, K.about.S.sub.0.2, so that the probe advantageously operates in the limit region K.about.0.1. The problem thus arises that changes in the reflection space which increase K lead to a noticeable reduction of the probe sensitivity. Such a modification occurs given blood or tissue deposits on the probe which unavoidably occur given intracardial implantation. The deposits reduce the sensitivity of the probe not only by damping the transition of the light between the sensor and blood and, thus, the factor (dk/dS.sub.02) but also by changing K as a consequence of the light reflection at the deposit which is not S.sub.O2- dependent, referred to below as "zero reflection".
The decrease in the sensitivity can also be explained because, in the known arrangement, the transmission current flowing through the light transmitter is reduced by the reception current flowing through the light receiver, thus that much less transmission current is available given increasing zero reflection.
FIG. 3 shows the standardized sensitivity E.sub.ni for constant current operation dependent on deposits having the thickness d at the measuring probe. The sensitivity E.sub.ni is thereby standardized such that the value I is without deposits. The curve III is valid for the prior art. The great decrease of the sensitivity with increasing deposits is thereby clearly visible. This decrease of the sensitivity creates problems with respect to the resolution range of the measuring amplifier. In order to keep the circuit outlay from becoming excessively high, measuring amplifiers having too high a dynamic range cannot be employed.